Acute heart failure is the sudden inability of the heart to fill with or pump a sufficient volume of blood. The afflicted patient may become weak and short of breath and, in some instances, die. In the most severe acute heart failure episodes, the patient may suffer from cardiogenic shock, a condition that is associated with high reported mortality rates.
Acute heart failure occurs in a variety of contexts. For example, some patients who are hospitalized for Acute Coronary Syndrome (i.e., heart attack and unstable angina) develop acute heart failure. Furthermore, some open-heart surgery patients develop acute heart failure. Acute heart failure also complicates certain illnesses. Additionally, some patients who undergo a Percutaneous Coronary Intervention or other procedure are at risk for developing acute heart failure or dying.
Acute heart failure does not necessarily progress to chronic heart failure or death; recovery is possible. Many patients who have acute heart failure and those at risk for developing it receive interventions that are intended to temporarily assist the heart during a recovery period. The intervention typically lasts for less than a week, but may last longer.
These interventions may include pharmaceuticals and medical devices, including cardiac-assist devices. When these cardiac-assist devices include a pump that supplements the heart's pumping action, they often are referred to as “blood pumps.” An effective cardiac assist device assumes some of the heart's pumping function, thereby unloading the heart and enabling it to recover. Cardiac-assist devices and blood pumps can be temporary or permanent.
The most common temporary cardiac-assist device is the intra-aortic balloon pump (“IABP”). An IABP is an inflatable balloon attached to a catheter. The IABP is inserted percutaneously (minimally invasively) into a peripheral vessel and advanced to the descending aorta. When the balloon inflates, it increases blood flow to the coronary arteries. When it deflates, it decreases the pressure against which the heart must pump. The IABP does not, however, increase the cardiac output significantly and a substantial percentage of cardiogenic shock patients that are implanted with IABPs die.
Other temporary cardiac-assist devices include extracorporeal (outside the body) blood pumps. Some of these devices require cardiac surgeons to connect them to the patient's heart and blood vessels directly through the exposed chest using “cannulas,” which are relatively large-diameter tubes. Such procedures are considered invasive, may require cardiopulmonary bypass, and are associated with significant complications. Some other extracorporeal blood pumps are connected to the patient using relatively wide catheters or cannulas that are inserted through peripheral blood vessels. Certain of these devices do not increase the heart's output significantly, are difficult to use, and/or are associated with significant complications.
Emerging data indicate that temporary, percutaneously-inserted blood pumps may provide an alternative to IABPs as well as extracorporeal blood pumps and other cumbersome devices. When these temporary blood pumps are attached to a catheter, they are known as “catheter blood pumps.” Some catheter blood pumps are inserted using established cath-lab techniques. These techniques are less invasive than cardiac surgery or other relatively complicated procedures.
Notwithstanding its attractiveness as a less-invasive alternative, certain percutaneously-inserted blood pumps exhibit one or more of the following drawbacks, in addition to any other shortcomings:
limited pump flow;
some degree of hemolysis (i.e., destruction of red blood cells);
require the use of a large catheter/cannula, with a risk of ischemia; and
relatively high cost.
The limited pump flow results from the fact that to be inserted percutaneously, the blood pump must be quite small. In particular, it is desirable for the blood pump to have a 12 French (4 millimeter) or smaller catheter.
Hemolysis can result when blood comes into contact with rapidly rotating elements.
One approach in the prior art to the size/flow challenge is the “expandable” blood pump. This pump, which is suitably small for percutaneous insertion, expands once in place within the vasculature or the heart. Although interesting conceptually, the expandable blood pump has proven to be difficult to implement. The pumps disclosed in U.S. Pat. Nos. 4,753,221, 5,749,855, and 6,981,942 are representative of the expandable blood pump and the problems of its implementation.
The pump that is disclosed in the '221 patent includes a catheter having an expandable propeller that is disposed in its distal end. When the propeller is deployed for operation, it spans a distance that is greater than the diameter of the catheter. To deploy the propeller, therefore, the distal end of the catheter must be enlarged. For that purpose, the distal end of the catheter is formed from a flexible material that is capable of expanding outward to provide a flared, enlarged-diameter region that can accommodate the propeller. The end of the catheter is enlarged by inflating a balloon that is disposed on the exterior of the catheter, just proximal to the flexible region. In particular, the tension on the distal end of the catheter resulting from the inflating balloon causes the enlargement.
Depending upon the orientation of the pump within the heart, blood is either: (1) drawn through the catheter and expelled at its distal end (near the pump), or (2) drawn in at the distal end of the catheter near the propeller, pumped through a length of the catheter, and then expelled through orifices. In either case, pumping the blood through the relatively smaller diameter catheter substantially defeats the purpose of providing an expandable pump. The region for flow must be expanded, as well as the propeller, to reap the benefits of increased flow.
The pump that is disclosed in the '855 patent has a drive cable that is surrounded, near its distal end, by an expandable cage. The distal end of the drive cable terminates in a spherical protuberance or “ball.” This ball is received in a socket that is formed in the distal end of the cage. This ball and socket arrangement serves as the distal bearing of the pump. The proximal end of the cage merges into a sleeve that surrounds the drive cable. In the absence of an applied, axially-directed force, the cage remains in a collapsed state. In this state, the cage has a cylindrical form that closely surrounds the drive cable and enables the pump to be inserted into a catheter.
An outwardly foldable propeller depends from the drive cable a short distance from its distal end. In the absence of an applied, axially-directed force, the propeller remains folded flat against the drive cable. The drive cable consists of an inner part (which extends distal to the propeller and terminates in the ball, as discussed above) and an outer part (which ends at the propeller). The drive cable is designed so that its inner part is movable relative to its outer part. As the inner part of the drive cable is drawn in the proximal direction by an axially-applied force (e.g., by a medical practitioner tugging on the inner part), relative movement between the inner and outer parts of the drive cable expands the propeller. At the same time, and by virtue of the same applied force, relative movement between the sleeve and the outer part of the drive cable expands the cage. The deployed propeller can then freely spin within the expanded cage.
While this pump overcomes the aforementioned flow-restriction problem with the '221 patent, it suffers from several other significant drawbacks. One drawback pertains to its distal bearing. This bearing, which stabilizes the propeller within the expandable cage, is implemented as a thrust bearing. That is, in operation of the pump, the bearing is placed in tension as the inner part of the drive cable is moved axially (and subsequently held in place) to expand the cage and propeller. It will be appreciated that under tension, this bearing is difficult to seal. And any blood that enters the space between the ball (which is rapidly spinning since the cable is spinning) and the socket will be hemolyzed and otherwise disrupted, making this bearing a likely site for thrombosis. Furthermore, the sliding friction between the ball and socket results in heat and wear. The heat can damage the blood and the wear can generate particulates.
A second shortcoming with the blood pump that is disclosed in the '855 patent relates to the fact that the drive cable experiences an axial load (to keep the cage expanded). This axial load increases the severity of wear between the drive cable and the surrounding catheter since both will be generally non-rectilinear (and thus establish loci of high surface contact pressure) to reach the heart from the insertion point at the femoral artery.
A third drawback of this blood pump is that it requires at least three seals: one for the distal bearing, a second for the two-part drive cable, and a third for the sleeve and outer part of the drive cable. Every seal presents a possibility for leakage of blood past the seal, which can ultimately result in hemolysis and thrombosis.
A third prior-art blood pump, as disclosed in the '942 patent, expands via inflation. In particular, the pump includes an inflatable propeller and a surrounding inflatable housing. As the housing inflates, it expands outwardly as well as inwardly. Since there appears to be nothing to restrict inward expansion of the housing, the inflated housing and the inflated propeller are likely to come into contact with one another. Contact between the rotating propeller and the housing could interfere with proper pumping function.
The problems with prior-art expandable blood pumps, as exemplified by the shortcomings of the three devices discussed above, limit their potential utility as a life-sustaining device. As a consequence, cardiac specialists and the patients that they treat would benefit from improvements to percutaneously-inserted, expandable blood pumps.